Poly(methylmethacrylate) (PMMA) has been widely used in orthopedic surgery as a bone cement and has long been known for its superior biocompatability. In a typical orthopedic implant procedure, PMMA cement is used as a grouting agent to fix the rigid stem (usually metal) of a prosthesis in the intramedullary canal of a bone such as, for example, the femur (as in total hip arthroplasty). PMMA bone cement conventionally includes an acrylic polymeric powder which is mixed in the operating room with a liquid acrylic monomer system to provide a dough mass. The doughy mass is inserted into the prepared intramedullary cavity and then, while the cementitious mixture is still in a semi-fluid state, the stem of the prosthesis is fitted into the canal. Within a few minutes, polymerization converts the semi-fluid grout into a hardened mantle.
Despite its advantages in terms of biocompatability, PMMA cement is relatively weak (compared to the bone and the implanted stem) and is frequently found unable to withstand the long-term cyclic loading experienced by a prosthetic joint. Over time, fatigue cracking of the cement mantle may occur along with breakdown of the metal-cement interface. Such fatigue cracking may progress to the point at which there is a loss of support of the metal stem in the canal, resulting in the device becoming loose, unstable and painful. The ultimate result may be a need for replacement of the prosthesis, a difficult and painful procedure.
Another limiting aspect of PMMA as an orthopedic cement material is that there are significant drawbacks to having such cement polymerize in the body cavity. Sufficient heat may evolve during the setting reaction to cause tissue damage and necrosis. Also, the monomer itself has been considered toxic and, if it diffuses from the polymerizing mass, local as well as systemic effects can result (including death). During polymerization, there is an associated shrinkage of the PMMA cement of as much as 10% and such shrinkage may cause residual stresses and premature failure of the mantle. Also, during insertion of the prosthetic stem into an intramedullary canal, it is difficult to assure an optimal mantle thickness of about 2 to 4 mm everywhere about the stem and, if substantial variations occur, the non-uniform thickness may accelerate fracture and fragmentation of the mantle.
Some of these problems are discussed in U.S. Pat. No. 4,491,987. In an effort to improve the interfacial bond between the stem of a prosthesis and the bone cement applied at the time of implantation, the patent teaches that the stem, preferably textured or manufactured with a porous outer surface, should be precoated with a thin layer of PMMA. Because of the precoating, a lesser amount of new bone cement is employed during the subsequent surgical procedure. The exotherm of the reaction is thus limited, decreasing the probability of necrosis and reducing the possibility of systemic interference resulting from toxic monomer.
While such a precoat enhances implant-cement interfacial strength by having the new cement bond to the PMMA precoat (instead of directly to the metal or ceramic stem) during implantation, the mechanical properties of the acrylic precoat in terms of strength, modulus, and fracture toughness are not notably superior to those of bulk acrylic.
Considerable effort has been expended to improve the properties of PMMA so that its fatigue behavior more closely matches that of a prostheses it fixes in place. Some of that effort has involved the reinforcement of PMMA with high strength fibers of stainless steel, carbon, or Kevlar. See B. M. Fishbane and S. R. Pond, Clin. Orthop. Rel. Res., Vol. 128, p. 194 (1977); R. M. Pilliar and R. Blackwell, J. Biomed. Mater. Res., Vol. 10, p. 893; S. Saha and S. Pal, Trans. 7th Ann. Soc. Biomater., Vol. 4, p. 21 (1981). However, the inclusion of such fibers in a composite bone cement tends to increase the viscosity of the semi-fluid mixture, making application more difficult and increasing the possibility that objectionable voids or windows may occur in the cement mantle. Also, the properties of these composites are controlled by the strength of the fiber-matrix bond which, for the fibers mentioned, is fairly low.
Whether such fibers are incorporated in the acrylic cement applied at the time of surgery or in a precoat applied to the stem of a prosthesis, they introduce an additional material that may create or complicate problems of biocompatability. Such concerns would be reduced if the reinforcing fibers in a PMMA cement matrix were of a like material.
A process for producing higher-strength PMMA fibers for possible use in reinforcing a PMMA matrix has been described in an article by C. A. Buckley, E. P. Lautenschlager and J. L. Gilbert in J. Applied Polymer Science, Vol. 44, pp. 1321-1330 (1992), the disclosure of which is incorporated by reference herein. In that process, PMMA was drawn into fibers by melt extrusion followed immediately by a transient temperature drawing process. By adjusting processing variables, fibers ranging from 25 .mu.m to 635 .mu.m in diameter were produced. Those fibers produced by a relatively slow extrusion speed and small extrusion hole diameter combined with a relatively fast draw rate were found to have the highest degree of molecular orientation or alignment as reflected by their relatively high heat relaxation ratios. Both tensile strength (Ultimate Tensile Strength) and modulus increased dramatically with greater molecular orientation, as reflected by length relaxation ratios. For example, a maximum UTS of 225 MPa (megapascals) was observed in a fiber of 36 .mu.m diameter having a length relaxation ratio of 18.7 to 1, representing approximately a 600% increase in strength over bulk PMMA material.
Other references indicating the state of the art are U.S. Pat. Nos. 4,963,151, 4,735,625, 5,037,442, 4,895,573, 3,992,725, 4,718,910, 4,851,004, 5,080,680, 5,180,395, 5,197,990, 4,743,257 and 5,171,288.